With the average lifespan and age of the population on the increase, vascular diseases are guaranteed to strike growing numbers within the population. Among such diseases are neurovascular disorders, which encompass those conditions that result in cerebrospinal ischemia, infarction, and hemorrhage. To provide an example, every year, over 700,000 people in the United States suffer a stroke, and roughly a quarter of those strokes are fatal. Stroke is therefore the third leading cause of death in the United States. In addition, a large segment of the elderly population is debilitated by dementia. Recently, neuronal vascular disorders, including microstrokes (lacunes), microbleeds, and neurovascular disease have been linked with many forms of dementia, such as Alzheimers's disease and vascular dementia. (Heye and Cervos-Navarro 1996; del Zoppo and Mabuchi 2003; Wardlaw, Sandercock et al. 2003) At present, options for treatment of stroke remain few and of limited efficacy despite years of basic and clinical research. Continued progress in stroke research depends critically on animal models that allow stroke to be studied at various stages, from initial changes in physiological parameters (e.g., blood flow and blood oxygenation) to neuronal death, behavioral impairment, and recovery. (del Zoppo 1998; Lipton 1999; del Zoppo and Mabuchi 2003). Most ischemic stroke models developed to date produce either large-scale injury, or a multitude of small-scale injuries at uncontrolled sites. Most hemorrhagic stroke models developed to date produce either large-scale hemorrhage or systemic injury. These existing models do not allow the production of small-scale, localized injury or blockage to specifically targeted vessels at depth. Such a paradigm is particularly crucial for the study of the effects of ischemic microstrokes and microbleeds.
Existing in vivo animal models of stroke fall into one of five broad categories: 1) occlusion of large vessels by ligation or filament insertion; 2) occlusion of a multitude of microvessels by injection of embolus into the bloodstream; 3) hemorrhagic damage (vessel rupture) by injection of a tissue-degrading substance; 4) model of hemorrhage by injection of whole or fractionated blood; and 5) optically-induced thrombosis of blood vessels by linear absorption of light. There is no reported technique that is capable of producing both thrombotic and hemorrhagic stroke to specific individual vessels deep within the same preparation.
In the case of mechanical occlusion, current techniques involve the blockage of blood flow by a variety of methods. These methods include ligation of large arteries (e.g., carotid artery) (McBean and Kelly 1998)], ligation of smaller arteries (Wei, Rovainen et al. 1995; Wei, Erinjeri et al. 2001), and insertion of a filament into a large artery for the occlusion of a main arterial branch. (e.g., the middle cerebral artery) (Tamura, Graham et al. 1981; Chen, Hsu et al. 1986; Busch, Kruger et al. 1998). Artery ligation results in neuronal injury to large, millimeter or larger sized regions of the rodent brain and is a model for major infarcts.
As a model for microstrokes, microspheres (Lyden and Hedges 1992; Lyden, Zivin et al. 1992; Lyden, Lonzo et al. 1997) or preformed clots (Kudo, Aoyama et al. 1982; Overgaard 1994; Krueger and Busch 2002) can be injected into an artery, leading to occlusion of smaller vessels downstream from the injection site, but without allowing specific individual vessels to be targeted. As a result, physiological changes cannot be correlated to specific local disruptions.
Hemorrhages can be induced by systemic or local injections of agents such as collagenase (Rosenberg, Mun-Bryce et al. 1990)], or tissue plasminogen activator (tPA) (Dijkhuizen, Asahi et al. 2002) to weaken vessels or disrupt the blood-brain barrier. Using such models to evaluate potential treatments is difficult because the effects of these agents can be spread over large, uncontrolled volumes. In addition, the effects of the agent cannot be isolated to the vasculature alone, because the agent can directly affect the surrounding tissue.
Alternatively, direct injection of whole or fractionated blood into the extracellular space has been reported as a model for hemorrhagic stroke (Deinsberger, Vogel et al. 1996; Hickenbottom, Grotta et al. 1999). The spatial localization is limited by diffusion of the injected materials. Additionally, this model is deficient in other aspects of natural hemorrhagic stroke, including the vascular and endothelial response. Current models of hemorrhage cannot be used as models of small hemorrhage, which are necessary for studies of vascular dementia.
For the case of optically-induced thrombosis, previous work utilized green light to excite an intravenously injected photosensitizer. When excited by exposure to light, photosensitizers generate singlet oxygen (Pooler and Valenzeno 1981), which attacks the membranes of the vessel walls (Herrmann, 1983). Damage to the vessel walls then starts a natural cascade of activation that results in the formation of a clot in all exposed vessels (Watson, Dietrich et al. 1985; Krammer 2001). In earlier work, transcranial illumination with diffuse green light exposed blood vessels over a wide lateral and axial extent, 1-3 millimeters in diameter (Watson, Dietrich et al. 1985; Dietrich, Ginsberg et al. 1986; Dietrich, Ginsberg et al. 1986).
More recently, work has been done using green light that is tightly focused through a microscope objective, constraining the lateral dimension of exposure at the focal plane to approximately one micrometer (Schaffer, Ebner et al. 2003; Schaffer, Ebner et al. 2003; Schaffer, Tsai et al. 2003), allowing individual vessels to be clotted. While very powerful, this focal photothrombotic stroke model has one major drawback: localized clotting can be achieved only in surface vessels. This limitation is due to the single-photon excitation of the photosensitizer molecule. When focused on a deep-lying vessel to induce a clot, all vasculature lying above that vessel is also clotted, preventing the use of this model for studying the effect of localized thrombosis in individual vessels at depth.
Alternatively, highly absorbed wavelengths of light (e.g., 10.2 microns from a CO2 laser) are used extensively in neurosurgery to simultaneously remove neuronal and vascular tissue while concurrently cauterizing the remaining portions of the removed blood vessels. The mechanism of damage relies on the linear absorption of the laser light by water and other tissue constituents and, therefore, does not require the presence of an exogenous photosensitizer. The high absorption coefficient of the tissue at these wavelengths results in substantial energy absorption and thermal buildup within the targeted tissue. The concurrent thermal diffusion out of the targeted volume results in an extended region of collateral thermal damage.
Current medical treatment for stroke requires therapeutic intervention within hours of the stroke to be optimally effective. Full understanding of the mechanisms and efficacy of these interventions therefore requires real-time visualization of stroke with high spatial and temporal resolution. Previously, real-time visualization and quantification of the effects of vascular damage on blood flow and blood vessel morphology have been performed using technologies, such as laser Doppler flowmetry (Dirnagl, Kaplan et al. 1989; Nakase, Kakizaki et al. 1995), magnetic resonance imaging (MRI) (Hoehn-Berlage, Norris et al. 1995; Busch, Kruger et al. 1998), positron emission tomography (PET) (Marchal, Young et al. 1999), computer-aided tomography (CAT), fluorescent video microscopy (Wei, Rovainen et al. 1995; Wei, Erinjeri et al. 2001; Ishikawa, Sekizuka et al. 2002), or confocal laser scanning microscopy (Seylaz, Charbonne et al. 1999; Pinard, Nallet et al. 2002). With the exception of the light microscopy, these techniques are limited to determining average blood flow over 100-1000-micrometer-sized areas. While such averages may be relevant for determining the degree of ischemia or hemorrhage, they provide no input on changes in flow and morphology in individual vessels, save for the largest branches of the cerebral vasculature. Fluorescent video microscopy allows individual vessels to be studied, but is limited to the observation of surface vessels only, while confocal microscopy allows vessels up to approximately 50 μm beneath the surface to be visualized. These observation techniques have allowed quantitative characterization of changes in blood flow velocity and blood vessel dilation as a result of large-scale ischemia produced by surgical occlusion of arteries and arterioles (Wei, Rovainen et al. 1995; Wei, Craven et al. 1998; Seylaz, Charbonne et al. 1999; Wei, Erinjeri et al. 2001; Pinard, Nallet et al. 2002). These studies could not, however, address local changes in blood flow and vessels near an isolated occlusion. Recently, fluorescent video microscopy was used to study vessel dilation after photochemically-induced clots in individual arterioles, but the results were limited to surface vessels and blood flow could not be resolved (Ishikawa, Sekizuka et al. 2002).
Another modality of analysis for current models of induced stroke is based on the observation of behavior deficits in the subject and post-mortem histology of the targeted and collateral tissue regions. These widely utilized methods are performed hours to days after the onset of damage and, therefore, are unable to elucidate the dynamics and mechanisms involved in the propagation of injury due to vascular damage.
The study of microstrokes and microhemorrhages requires microscopic resolution, coupled with the ability to either precisely target or locate the microscopic vascular disturbance within the brain volume. Using nonlinear microscopy, local changes in blood flow due to isolated occlusions can be studied and quantified in real-time.
The use of nonlinear optical effects to provide contrast for image formation has revolutionized microscopy over the past decade. Many nonlinear effects are now used for imaging, including second- and third-harmonic generation, Coherent Anti-Stokes Raman scattering, the Kerr effect, and multi-photon excited fluorescence.
One non-linear technique is two-photon laser scanning microscopy, or “TPLSM” (Denk, Strickler et al. 1990; Denk 1994), which allows fluorescence imaging with intrinsic optical sectioning deep inside scattering specimens with diffraction-limited resolution. Briefly, an ultrashort laser pulse is tightly focused inside a specimen tagged with a fluorescent molecule that does not linearly absorb at the wavelength of the ultrashort laser. At the laser focus, the laser intensity can become high enough to induce two-photon excitation of the fluorescent molecule. Because the excitation is nonlinear, this fluorescence is only produced in the focal volume where the laser intensity is high. The fluorescence intensity is then recorded as the position of the laser focus is scanned throughout the specimen forming a three-dimensional image. In addition, because photoexcitation occurs only at the laser focus, there is significantly reduced bleaching of fluorescent dyes and photodamage to the sample as compared to linear imaging techniques.
TPLSM is especially well suited to in vivo imaging deep into highly scattering specimens, such as brain. In widefield or confocal fluorescence microscopy, the fluorescence must be imaged to a camera or to a pinhole, respectively. Scattering of the fluorescence leads to an unwanted background in widefield microscopy and to decreased signal strength in confocal microscopy. In TPLSM, however, because all the fluorescence originates from the focal volume, it need only be detected in order to contribute to the signal, not imaged. Thus fluorescence that is scattered on the way to the detector still contributes to image formation, and does not produce unwanted background. This immunity to scattering of the fluorescence allows imaging deep into scattering samples. The imaging depth is ultimately limited by scattering of the ultrashort laser beam. In practice, one can image up to 500 micrometer beneath the cortical surface in rat (providing access to layers 1-4 of the cortex), without loss of image resolution. For neuronal tissue with labeling throughout the tissue, a theoretical limit for imaging depth is approximately 1 millimeter.
TPLSM further provides means for measuring and quantifying the velocity, i.e., direction and speed, and flux of red blood cell (RBC) movement and plasma flow in vivo under acute as well as chronic conditions. These measurements make use of either fluorescently labeled plasma, in which case cells in the blood, such as RBCs and leukocytes, appear as dark objects on a bright background, or the use of fluorescently labeled RBCs.
The above-described technologies provide means for forming and observing strokes. However, these techniques for induction of stroke are incapable of producing hemorrhage, thrombosis, and breach of the blood-brain barrier targeted to specific individual blood vessels. Further, no technique is currently available for the production of surface or subsurface vascular injury localized with micrometer precision, thereby permitting the disruption of the smallest vessels, i.e., capillaries. Accordingly, the need remains for a device and method with such capabilities.